The present invention relates to a data reducer for a scintillation camera.
In the human body, increased metabolic activity is associated with an increase in emitted radiation. In the field of nuclear medicine, increased metabolic activity within a patient is detected using a radiation detector such as a scintillation camera.
Scintillation cameras are well known in the art, and are used for medical diagnostics. A patient ingests, or inhales or is injected with a small quantity of a radioactive isotope. The radioactive isotope emits photons that are detected by a scintillation medium in the scintillation camera. The scintillation medium is commonly a sodium iodide crystal, BGO or other. The scintillation medium emits a small flash or scintillation of light, in response to stimulating radiation, such as from a patient. The intensity of the scintillation of light is proportional to the energy of the stimulating photon, such as a gamma photon. Note that the relationship between the intensity of the scintillation of light and the gamma photon is not linear.
A conventional scintillation camera such as a gamma camera includes a detector which converts into electrical signals gamma rays emitted from a patient after radioisotope has been administered to the patient. The detector includes a scintillator and photomultiplier tubes. The gamma rays are directed to the scintillator which absorbs the radiation and produces, in response, a very small flash of light. An array of photodetectors, which are placed in optical communication with the scintillation crystal, converts these flashes into electrical signals which are subsequently processed. The processing enables the camera to produce an image of the distribution of the radioisotope within the patient.
Gamma radiation is emitted in all directions and it is necessary to collimate the radiation before the radiation impinges on the crystal scintillator. This is accomplished by a collimator which is a sheet of absorbing material, usually lead, perforated by relatively narrow channels. The collimator is detachably secured to the detector head, allowing the collimator to be changed to enable the detector head to be used with the different energies of isotope to suit particular characteristics of the patient study. A collimator may vary considerably in weight to match the isotope or study type.
Scintillation cameras are used to take four basic types of pictures: spot views, whole body views, partial whole body views, SPECT views, and whole body SPECT views.
A spot view is an image of a part of a patient. The area of the spot view is less than or equal to the size of the field of view of the gamma camera. In order to be able to achieve a full range of spot views, a gamma camera must be positionable at any location relative to a patient.
One type of whole body view is a series of spot views fitted together such that the whole body of the patient may be viewed at one time. Another type of whole body view is a continuous scan of the whole body of the patient. A partial whole body view is simply a whole body view that covers only part of the body of the patient. In order to be able to achieve a whole body view, a gamma camera must be positionable at any location relative to a patient in an automated sequence of views.
The acronym xe2x80x9cSPECTxe2x80x9d stands for single photon emission computerized tomography. A SPECT view is a series of slice-like images of the patient. The slice-like images are often, but not necessarily, transversely oriented with respect to the patient. Each slice-like image is made up of multiple views taken at different angles around the patient, the data from the various views being combined to form the slice-like image. In order to be able to achieve a SPECT view, a scintillation camera must be rotatable around a patient, with the direction of the detector head of the scintillation camera pointing in a series of known and precise directions such that reprojection of the data can be accurately undertaken.
A whole body SPECT view is a series of parallel slice-like transverse images of a patient. Typically, a whole body SPECT view consists of sixty four spaced apart SPECT views. A whole body SPECT view results from the simultaneous generation of whole body and SPECT image data. In order to be able to achieve a whole body SPECT view, a scintillation camera must be rotatable around a patient, with the direction of the detector head of the scintillation camera pointing in a series of known and precise directions such that reprojection of the data can be accurately undertaken.
Therefore, in order that the radiation detector be capable of achieving the above four basic views, the support structure for the radiation detector must be capable of positioning the radiation detector in any position relative to the patient. Furthermore, the support structure must be capable of moving the radiation detector relative to the patient in a controlled manner along any path.
In order to operate a scintillation camera as described above, the patient should be supported horizontally on a patient support or stretcher.
The detector head of the scintillation camera must be able to pass underneath the patient. Therefore, in order for the scintillation camera to generate images from underneath the patient, the patient support must be thin. However, detector heads are generally supported by a pair of arms which extend from a gantry. Thus, the patient support generally must be cantilevered in order for the detector head to be able to pass underneath the patient without contacting any supporting structure associated with the patient support. The design of a cantilevered patient support that is thin enough to work properly with a scintillation camera is exceedingly difficult. Expensive materials and materials that are difficult to work with, such as carbon fibre, are often used in the design of such cantilevered patient supports.
A certain design of gantry or support structure for a scintillation camera includes a frame upon which a vertically oriented annular support rotates. Extending out from the rotating support is an elongate support. The elongate generally comprises a pair of arms. The pair of arms generally extends through a corresponding pair of apertures in the rotating support. One end of the pair of arms supports the detector head on one side of the annular support. The other end of the pair of arms supports a counter balance weight. Thus, the elongate support is counterbalanced with a counterweight on the opposite side of the detector head.
With such a design of support structure for a scintillation camera, a patient must lie on a horizontally oriented patient support. The patient support must be cantilevered so that the detector head can pass underneath the patient. If the detector head must pass underneath only one end of the patient, such as the patient""s head, the cantilevered portion of the patient support is not long enough to cause serious difficulties in the design of the cantilevered patient support. However, if the camera must be able to pass under the entire length of the patient, the entire patient must be supported by the cantilevered portion of the patient support. As the cantilevered portion of the patient support must be thin so as not to interfere with the generation of images by the scintillation camera, serious design difficulties are encountered.
Among the advantages associated with such as design of support structure is that a patient may be partially pass through the orifice defined by the annular support so that the pair of arms need not be as long. However, the patient support must be able to support the patient in this position relative to the annular support, must be accurately positionable relative to the annular support, and must not interfere either with the rotation of the annular support or with the cables which will inevitably extend from the detector head to a nearby computer or other user control.
The photomultiplier tubes in a scintillation camera generate electric signals. The signals are processed, and images are created corresponding to the radiation emitted by the patient.
U.S. Pat. No. 3,011,057 issued to H. O. Anger, discloses a typical scintillation camera of the type described, wherein the photomultiplier are arranged in a hexagonal pattern over a circular crystal, and have overlapping fields of view. A hexagonal pattern is selected because it achieves the densest clustering possible of photomultiplier, having circular or hexagonal photocathodes.
Computation of the displacement of a light event from each of the two orthogonal coordinate axes is achieved by weighting the outputs of each of the photomultipliers in accordance with its distance from the coordinate axis in question, and summing the outputs of photomultipliers. The weighted sum of the photomultiplier signals used to calculate the displacement of a light event from a coordinate axis represent a fixed analytical function of the signals. Because a single analytical function is used for computation purpose irrespective of the location in the crystal of a light event, the two parameters that are measures of the quality of performance of a scintillation camera of the type described, namely spatial resolution and uniformity, are spatially dependent (i.e., are dependent on the location of the light event in the crystal). In other words, the resolution and uniformity for a given analytical function of the photomultiplier signals may be much better for events that occur in one region of the crystal as compared to events that occur in other regions.
Another conventional camera of the type described is disclosed in U.S. Pat. No. 3,171,763 issued to Tanaka et al. In this camera, the coordinate position of a photomultiplier establishes a delay time by which the photomultiplier signals can be separated in time sequence. The maximum resolution and linearity of this camera depend on the similarity between the shape of the electronic pulse and the shape of the waveform produced as a result of geometric configuration of the device. Thus, this detector used time domain as the basis for calculating position and, as a consequence, has a relatively long dead time.
Another conventional camera of this type described is disclosed in U.S. Pat. No. 4,060,730 issued to Zioni et al. This describes a camera in which summed row and column data is used as the basis for computing circuitry coupled via an ADC to the photomultiplier to compute the projection of a light event in the crystal on a reference axis by forming an analytical function of the signals of the photomultipliers according to the spatial location of the light event in the crystal.
One problem to overcome in the design of scintillation cameras is that the computers generally used are not able to process the entirety of the data at the rate that it is generated by the photomultiplier tubes and associated electronic circuitry. It is necessary, therefore, to implement a method to reduce the quantity of data processed by the scintillation camera""s computer. In other words, only the useful values generated by the photomultiplier tubes must be selected, that is, the rest of the values are noise and therefore not useful.
Referring to FIG. 13, the signal generated by a photomultiplier tube rises quickly, that is, in about 200 nanoseconds, and decays more slowly, that is, in about 1 microsecond. It is in this short time period that the light signal must be collected. How effectively light is collected during this period is important, because the performance of a scintillation camera is directly related to how completely light is collected. A camera that misses too much light will generally have poor performance. Thus, it is important that the signal selection process does not interfere with the collection of light signal.
One light collection method known in the art is the threshold method. The threshold method is typically set up by assigning to the output an offset below zero volts. Thus, when the signal rises above zero volts, acceptance of the signal begins. However, one disadvantage is that when the signal drops below zero volts again, the signal is no longer accepted. This means that the integral of the signal between the offset voltage and zero volts is lost. The threshold method is slow, has an inherent bandwidth limitation, and distorts the signal.
An object of the invention is to provide an improved data reducer for a scintillation camera.
A second object of the invention is to provide a data reducer for a scintillation camera that effectively selects signals from photomultiplier tubes while minimizing distortion of the signals. According to one aspect of the present invention, there is provided a method of reducing data in the localization processing of scintillation events in a scintillation camera having a plurality of photomultiplier tubes arranged in rows and columns. The method comprises steps of: summing output signals from the photomultiplier tubes of each row to provide a summed row signal for each row; summing output signals from the photomultiplier tubes of each column to provide a summed column signal for each column; selecting a largest summed row signal and a largest summed column signal from the summed row and column signals; and determining at least three output signals to be required for computing the location of a scintillation event by using the largest summed row and summed column signals, thereby reducing the amount of data to be processed. The determining step can comprise steps of: locating an X and Y coordinate corresponding to the largest summed row and summed column signals; and choosing the output signals from at least three photomultiplier tubes which surround the X and Y coordinate.
In an embodiment of the invention a data reducer for a scintillation camera is applied when photomultiplier tubes in an array thereof generate output signals following a flash of light from a scintillator. The output signal from a given photomultiplier tube is connected to an integrating preamplifier circuit. Each output signal is connected to an analog to digital converter, to a row amplifier circuit and to a column amplifier circuit. The signal from the analog to digital converter is provided via a bus to a data processor. The output signal of the data processor is set according to operational requirements, and then provided to a digital to analog converter. The signal from the row amplifier circuit is provided to a row summing circuit. The signal from the column amplifier circuit is provided to a column summing circuit. The respective signals from the row summing circuit and the column summing circuit are provided to an energy analyser. The energy analyser ensures that the signal is the result of a gamma event. The output of the digital to analog converter is compared with the output of the energy analyser. If the signal from the energy analyser conforms with the requirements of the digital to analog converter, then a valid signal is sent to the photomultiplier tube map and address. From the row summing circuit the largest row signal is sent to the photomultiplier tube map and address, and from the column summing circuit the largest column signal is sent to the photomultiplier tube map and address.
Advantageously, the invention provides: an improved data reducer for a scintillation camera; and a data reducer for a scintillation camera that effectively selects signals from photomultiplier tubes while minimizing distortion of the signals.
Other advantages, objects and features of the present invention will be readily apparent to those skilled in the art from a review of the following detailed description of preferred embodiments in conjunction with the accompanying drawings and claims.